This invention relates generally to detectors for computed tomography (CT) imaging systems, and more particularly to optimizations of such detectors for medical and other applications and to imaging systems using such optimized detectors.
In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the xe2x80x9cimaging planexe2x80x9d. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a xe2x80x9cviewxe2x80x9d. A xe2x80x9cscanxe2x80x9d of the object comprises a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector.
In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection technique. This process converts the attenuation measurements from a scan into integers called xe2x80x9cCT numbersxe2x80x9d or xe2x80x9cHounsfield unitsxe2x80x9d, which are used to control the brightness of a corresponding pixel on a cathode ray tube display. In another mode of operation of the CT imaging system, a helical scan is used to obtain projection data for images.
More particularly, and referring to FIGS. 1 and 2, one known computed tomograph (CT) imaging system embodiment 10 includes a gantry 12 representative of a xe2x80x9cthird generationxe2x80x9d CT scanner. Gantry 12 has an x-ray source 14 that projects a beam of x-rays 16 toward a detector array 18 on the opposite side of gantry 12. Detector array 18 is formed by detector elements 20 which together sense the projected x-rays that pass through an object 22, for example a medical patient. In at least one embodiment of the present invention, detector array 18 is fabricated in a multi-slice configuration. Each detector element 20 produces an electrical signal that represents the intensity of an impinging x-ray beam. As the x-ray beam passes through a patient 22, the bean is attenuated. During a scan to acquire x-ray projection data, gantry 12 and the components mounted thereon rotate about a center of rotation 24.
Rotation of gantry 12 and the operation of x-ray source 14 are governed by a control mechanism 26 of CT system 10. Control mechanism 26 includes an x-ray controller 28 that provides power and timing signals to x-ray source 14 and a gantry motor controller 30 that controls the rotational speed and position of gantry 12. A data acquisition system (DAS) 32 in control mechanism 26 samples analog data from detector elements 20 and converts the data to digital signals for subsequent processing. An image reconstructor 34 receives sampled and digitized x-ray data from DAS 32 and performs high speed image reconstruction. The reconstructed image is applied as an input to a computer 36 which stores the image in a mass storage device 38.
Computer 36 also receives commands and scanning parameters from an operator via console 40 that has a keyboard. An associated cathode ray tube display 42 allows the operator to observe the reconstructed image and other data from computer 36. The operator supplied commands and parameters are used by computer 36 to provide control signals and information to DAS 32, x-ray controller 28 and gantry motor controller 30. In addition, computer 36 operates a table motor controller 44 which controls a motorized table 46 to position patient 22 in gantry 12. Particularly, table 46 moves portions of patient 22 through gantry opening 48.
Multiple slice detector arrays 18 increase the rate at which a scan of a given volume can be performed by acquiring data for several parallel image slices at the same time. For example, and referring to FIGS. 3 and 4, one known prior art detector array 18 includes a plurality of detector modules 50. Each detector module includes an array of detector elements 20. Particularly, each x-ray detector module 50 includes a plurality of scintillators 52 positioned above and adjacent corresponding photodiodes 54, a semiconductor device 56, and at least one flexible electrical cable 58. Photodiodes 54 are either individual photodiodes or a multi-dimensional photodiode array. Photodiodes 54 are optically coupled to scintillators 52 and generate electrical outputs on lines 60, wherein the outputs are representative of light output by corresponding scintillators 52. Each photodiode 54 produces a separate electrical output 60 that is a measurement of the beam attenuation for a specific element 20. Photodiode output lines 60 are, for example, physically located on one side of module 50 or on a plurality of sides of module 50. As shown in FIG. 4, photodiode outputs 60 are located at top and bottom of the photodiode array.
Semiconductor device 56 includes two semiconductor switches 62 and 64. Switches 62 and 64 each include a plurality of field effect transistors (FET) (not shown) arranged as a multidimensional array. Each FET includes an input line electrically connected to a photodiode output 60, an output line, and a control line (not shown). FET output and control lines are electrically connected to flexible cable 58. Particularly, one-half of photodiode output lines 60 are electrically connected to each FET input line of switch 62 with the remaining one-half of photodiode output lines 60 electrically connected to the FET input lines of switch 64.
Flexible electrical cable 58 includes a plurality of electrical wires 66 connecting its ends. FET output and control lines are electrically connected to cable 58. Particularly, each FET output and control line is wire bonded to a wire 66 of one end of cable 58. FET output and control lines are wire bonded to wires 66 in the same manner as photodiode outputs (not shown) are wire bonded to the FET input lines (also not shown). Cables 58 are secured to detector module 50 using mounting brackets 68 and 70.
Referring to FIG. 5, after mounting detector modules 50 into detector array 18, unconnected cable 58 ends are coupled to DAS 32 so that an electrical path exists between photodiode 52 outputs and DAS 32, and so that FET control lines 72 are electrically connected to DAS 32 to enable semiconductor device FETs 74. In a four-slice CT imaging system 10 using the prior art detector array 18 embodiment of FIGS. 3, 4, and 5, each column of detector module 50 is electrically connected to four DAS 32 channels, i.e., two channels within each flexible electrical cable 58. (In general, an N channel system would have N channels connected to each column of detector module 50, with N/2 channels within each flexible electrical cable 58.) One exemplary channel is represented, in part, in FIG. 5. DAS 32 is coupled across a rotating gantry 12 slip ring 76 to computer 36 and image reconstructor or processor 34. Each detector element 20 includes a photodiode 54 that is coupled to a plurality of FETs 74, only one of which is shown. In a four-slice CT imaging system, each channel is coupled to the output of one-fifth of FETs 74. (Of the FETs not shown in FIG. 5, one set connects unused diode elements to ground during a scan.) Computer 36 provides a control signal instructing a controller 78 to turn on one or more FETs 74 per channel per data interval during an imaging scan, resulting in an analog signal from a corresponding one or more photodiodes 54 being applied to a preamp 82. The output signal from preamp 82 is converted to a digital signal by analog to digital converter 84 and sent across slip ring 76 to image reconstructor 34.
For reconstruction of medical images without motion artifacts, it is desirable to rotate gantry 12 as rapidly as possible to obtain a set of views for image reconstruction. It is correspondingly desirable to sample the outputs of photodiodes 54 as rapidly as possible to obtain images with as high a resolution as possible. However, the highest sampling rate is limited by the bandwidth of data communication across slip ring 76, among other things. In some applications, it is desirable to image as large an extent in the z-direction as possible in as little time as possible. For these applications, it has been necessary to effectively combine outputs of detector elements 20 in adjacent rows of detector array 18 transverse to the z-direction by turning on more than one FET 74 at a time. This combination allows a greater extent of a patient to be imaged in the z-direction in a shorter time, but the reconstructed images correspond to thicker slices of the imaging volume in the z-direction (i.e., lower z-axis resolution).
Detector elements 20 are only 1.25 mm in extent in the z-direction in one known detector array 18. Moreover, even though one known detector array 18 provides 16 rows of detector elements 20, one known imaging system 10 using such a detector array only provides sufficient DAS 32 electronics to process four image slices at a time. Therefore, cardiac imaging applications require either that a helical scan be performed or that multiple axial scans be performed, with table 46 being stepped between the axial scans. Providing more rows of detector elements 20 in detector modules 50 of detector array 18 would reduce the time needed to acquire data for a complete image of a patient""s heart, but this advantage could be gained only at the expense of a much greater number of DAS 32 channels.
It would therefore be desirable to provide a multislice detector array optimized for one or more imaging applications, including medical imaging applications. It would also be desirable to provide an imaging system using such a detector array that had a reduced need for additional DAS channels and additional bandwidth.
There is therefore provided, in one embodiment of the present invention, a detector array for a computed tomographic imaging system having a z-direction corresponding to an image slice thickness direction and that is arc-shaped in a direction transverse to the z-direction. The detector array has a plurality of detector modules configured so that the detector array has active regions of differing thicknesses.
This detector array embodiment provides an optimized detector array for certain imaging situations, for example, in cardiac imaging applications in which increased coverage is required only in a relatively small central portion of a field of view. Such detector array embodiments also reduce the number of detector acquisition system (DAS) channels and the corresponding bandwidth needed to process information from the detector array, because detector elements and their associated electronics are not provided where they are not needed.